A large variety of medical endoprostheses or implants for various applications are known from the prior art. Understood as implants within the meaning of the present invention are endovascular prostheses or other endoprostheses, for example stents, in particular wire mesh stents, attachment elements for bones, for example screws, plates, or pins, intramedullary pins, spiral bundle nails, Kirschner wires, wires for septal occluders, surgical suture material, intestinal clamps, vessel clips, prostheses for hard and soft tissue, and anchoring elements for electrodes, in particular for pacemakers or defibrillators.
Stents for the treatment of stenoses (vascular constrictions) are used particularly frequently as implants at the present time. Stents have a body in the form of an optionally perforated tubular or hollow cylindrical base lattice which is open at both longitudinal ends. The tubular base lattice of such an endoprosthesis is inserted into the vessel to be treated, and is used to support the vessel. Stents have become established in particular for the treatment of vascular diseases. Use of stents allows constricted regions in the blood vessels to be expanded, resulting in lumen gain. Although the optimal vessel cross section primarily necessary for successful treatment may be achieved by the use of stents or other implants, the permanent presence of such a foreign body initiates a cascade of microbiological processes which may lead to gradual overgrowth of the stent, and in the worst case may result in vascular occlusion. One approach to this problem is to fabricate the stent or other implants from a biodegradable material.
The term “biodegradation” refers to hydrolytic, enzymatic, and other metabolic degradation processes in the living organism which are primarily caused by the bodily fluids which come into contact with the biodegradable material of the implant, resulting in gradual disintegration of the structures of the implant containing the biodegradable material. As a result of this process, at a certain point in time the implant loses its mechanical integrity. The term “biocorrosion” is frequently used synonymously for “biodegradation.” The term “bioabsorption” includes the subsequent absorption of the degradation products by the living organism.
Suitable materials for the body of biodegradable implants may contain polymers or metals, for example. The body may be composed of several of these materials. The common feature of these materials is their biodegradability. Examples of suitable polymeric compounds include polymers from the group including cellulose, collagen, albumin, casein, polysaccharides (PSAC), polylactide (PLA), poly-L-lactide (PLLA), polyglycol (PGA), poly-D,L-lactide-co-glycolide (PDLLA-PGA), polyhydroxybutyric acid (PHB), polyhydroxyvaleric acid (PHV), polyalkyl carbonates, polyortho esters, polyethylene terephtalate (PET), polymalonic acid (PML), polyanhydrides, polyphosphazenes, polyamino acids, and the copolymers thereof, as well as hyaluronic acid. Depending on the desired characteristics, the polymers may be present in pure form, derivatized form, in the form of blends, or as copolymers. Metallic biodegradable materials are primarily based on alloys of magnesium and iron. The present invention preferably relates to implants whose biodegradable material contains, at least in part, a metal, preferably iron, manganese, zinc, and/or tungsten, in particular an iron-based alloy (referred to below as “iron alloy” for short).
In the implementation of biodegradable implants, the aim is to control the degradability corresponding to the intended treatment or use of the particular implant (coronary, intracranial, renal, etc.). For many therapeutic applications, for example, it is an important target corridor for the implant to lose its integrity over a period of four weeks to six months. In this regard “integrity,” i.e., mechanical integrity, refers to the characteristic that the implant does not undergo hardly any mechanical losses compared to the undegraded implant. This means that the implant is still mechanically stable enough to ensure that, for example, the collapse pressure drops only slightly, i.e., to a maximum of 80% of the nominal value. Thus, when integrity is present the implant is still able to fulfill its primary function of keeping the blood vessel open. Alternatively, integrity may be defined such that the implant is mechanically stable enough that in a load state in the blood vessel it to undergoes minimal changes in its geometry, for example does not show appreciable collapse, i.e., under a load of at least 80% of the dilation diameter, or, in the case of a stent, has very little fracturing of supporting struts.
Implants containing an iron alloy, in particular iron-containing stents, are particularly economical and easy to manufacture. For the treatment of stenoses, for example, these implants do not lose their mechanical integrity or support effect until after a comparatively long time, i.e., after a residence time of approximately two years in the treated organism. This means that for implants containing iron, the collapse pressure decreases too slowly over time for this application.
A long residence time of implants may also lead to complications if, as a result of its ferromagnetic properties, the implant undergoing disintegration does not permit examination of the treated patient using magnetic resonance tomography. Furthermore, for orthopedic implants (bone plates, for example), the formation of new bone substance may result in mechanical stresses between the implant which is degrading too slowly and the bone substance. This produces bone deformations or malformations, especially in children and adolescents. For such applications, therefore, it is desirable to be able to expand the potential of use in particular of iron-based implants through more rapid degradation.
Various mechanisms for controlling the degradation of implants have been described in the prior art. These mechanisms are based, for example, on inorganic and organic protective layers or a combination thereof which resist the human corrosive environment and the corrosion processes occurring therein. Previously known approaches are characterized by the achievement of barrier layer effects which are based on spatially separating, with as few defects as possible, the corrosion medium and a metallic material. As a result, the degradation time is increased. This ensures degradation protection by use of protective layers of various compositions and by defined geometric distances (diffusion barriers) between the corrosion medium and the magnesium base material. Other approaches are based on alloy components of the biodegradable material of the implant body which influence the corrosion process by shifting the position in the electromotive series. Further approaches in the area of controlled degradation produce predetermined breaking effects to by applying physical changes (for example, localized changes in cross section) in the stent surface (for example, local multilayers of chemically different compositions). However, the previously described approaches are usually not able to place the disintegration occurring due to the degradation process and the resulting strut fractures in the required time window. The result is degradation of the implant which begins either too early or too late, i.e., which has excessive variability.
Implants are known from documents EP 923 389 B1 (WO 99/03515 A2) or WO 2007/124230 A1 which are degradable in vivo as the result of corrosion. The material of the known implants contains iron as the primary component, as well as carbon in a specific, predetermined concentration. The disadvantage of these alloys is that the pure binary system composed of carbon and iron experiences a great loss in ductility with increasing carbon content, without a corresponding decrease in corrosion resistance.
In EP 08172105.2 various iron alloys, for example an alloy system based on Fe and Mn, are described which, although they are slightly more susceptible to corrosion compared to pure iron in an artificial medium which duplicates physiological conditions, they consistently degrade too slowly. A further disadvantage is that Mn stabilizes the cubic face-centered austenitic phase of the iron. This results in a mixed structure of austenitic and cubic body-centered ferritic phases, depending on the content of additional alloy elements and the preceding heat treatment, as well as the specific content of Mn itself. The particular percentage of these phases may fluctuate greatly from charge to charge on account of these numerous influencing variables. The resulting variations in hardening behavior of cut stents makes dimensioning, which is based on the assumption of specific strength and elongation values, difficult or impossible.